Effects of Changes in Ankle Joint Angle on the Relation Between Plantarflexion Torque and EMG Magnitude in Major Plantar Flexors of Male Chronic Stroke Survivors

I may be missing something but the results as listed have ABSOLUTELY NOTHING THAT WILL HELP RECOVERY.  Please read at the very bottom the Discussion and you will see what I mean; useless.

Effects of Changes in Ankle Joint Angle on the Relation Between Plantarflexion Torque and EMG Magnitude in Major Plantar Flexors of Male Chronic Stroke Survivors

Jongsang Son^1,2* and William Zev Rymer^1,2

• ¹Shirley Ryan AbilityLab (formerly the Rehabilitation Institute of Chicago), Chicago, IL, United States
• ²Department of Physical Medicine and Rehabilitation, Northwestern University, Chicago, IL, United States

The slope of the EMG-torque relation is potentially useful as a parameter related to muscular contraction efficiency, as a greater EMG-torque slope has often been reported in stroke-impaired muscles, compared to intact muscles. One major barrier limiting the use of this parameter on a routine basis is that we do not know how the EMG-torque slope is affected by changing joint angles. Thus, the primary purpose of this study is to characterize the EMG-torque relations of triceps surae muscles at different ankle joint angles in both paretic and non-paretic limbs of chronic hemispheric stroke survivors. Nine male chronic stroke survivors were asked to perform isometric plantarflexion contractions at different contraction intensities and at five different ankle joint angles, ranging from maximum plantarflexion to maximum dorsiflexion. Our results showed that the greater slope of the EMG-torque relations was found on the paretic side compared to the non-paretic side at comparable ankle joint angles. The EMG-torque slope increased as the ankle became plantarflexed on both sides, but an increment of the EMG-torque slope (i.e., the coefficient a) was significantly greater on the paretic side. Moreover, the relative (non-paretic/paretic) coefficient a was also strongly correlated with the relative (paretic/non-paretic) maximum ankle plantarflexion torque and with shear wave speed in the medial gastrocnemius muscle. Conversely, the relative coefficient a was not well-correlated with the relative muscle thickness. Our findings suggest that muscular contraction efficiency is affected by hemispheric stroke, but in an angle-dependent and non-uniform manner. These findings may allow us to explore the relative contributions of neural factors and muscular changes to voluntary force generating-capacity after stroke.

Introduction

Weakness of voluntary muscle contraction is a dominant clinical feature after hemispheric stroke, and the severity of such weakness is correlated with a stroke survivor's independence in performing many functional tasks (1). This reduction in voluntary muscle strength is typically the most obvious motor deficit (2), and many clinical phenomena observed in chronic stroke survivors could be attributed to a loss of strength rather than to a loss of control (3).

In addition to muscle weakness, muscular contractions in stroke-impaired muscles often appear less efficient than in contralateral muscles or in analogous muscles of intact subjects, potentially contributing to impaired voluntary force generation. To illustrate this assertion further (4), reported that in approximately half of the tested human stroke survivors, during sustained isometric voluntary contractions at different intensities, the slope of the biceps brachii (BB) electromyogram (EMG)-force relations was significantly greater on the paretic side than on the non-paretic side. A similar finding was also observed in paretic first dorsal interosseous (FDI) muscles of chronic stroke survivors (5), implying that paretic muscles may require the recruitment of more motor units to achieve a given muscle force, and thus, display inefficient muscular contractions.

It is evident then, that altered motor unit behavior may result in inefficient muscular contractions. For example, abnormally low mean motor unit firing rates were observed in paretic tibialis anterior (TA) muscles (6, 7) and in paretic BB muscles (8). There is also evidence for additional altered motor unit behavior, including compressed motor unit recruitment threshold ranges during isometric contractions of the paretic BB muscles (8), impairments in firing rate modulation in paretic BB muscles (9), disorganization in the rank order of recruitment in paretic FDI muscles (10), and a compressed range of motor unit firing rates in paretic FDI muscles (11). Moreover, a recent simulation study revealed that the recruitment compression and a compressed “onion-skin” firing pattern can potentially also contribute to voluntary muscle weakness (12). In addition, the slope of EMG-force relation was reported to increase in concert with compressed motor unit recruitment threshold ranges and reductions of mean motor unit firing rates (13). Although these earlier studies support the idea that altered neural factors can contribute, these neural factors may not be enough to explain a complicated EMG-force relation, because the EMG-force relation is an outcome of both neural and muscular factors.

In particular, muscular changes may also contribute to voluntary muscle weakness (14–17). Although our understanding of muscular changes after hemispheric stroke is still limited, it is relevant to note that decreased muscle thickness is often associated with reduced force output at a given level of muscle activation. In addition, the maximum isometric muscle force varies depending on the muscle length (18), and this force-length relation may also be substantially altered after stroke. For example, the width of the active force-length curve seems narrower in the paretic medial gastrocnemius (MG) muscles (19) than in the equivalent contralateral muscles. Such altered contractile properties may lead to modified torque-angle curves, showing a significant reduction in the magnitude of the torque at joint angles where muscle length is short (3, 20–22), potentially resulting in a higher slope of the EMG-torque relation at such a short length. This outcome is likely because the effective torque at shorter lengths is smaller at a given EMG. Furthermore, material properties of muscle tissues seem to contribute to muscle mechanics, revealing that a stiffer material surrounding contractile elements of muscle tissues can reduce fascicle strain as well as muscle force (23–26). Based on these observations, it is likely that muscular factors can also play a part in the abnormal EMG-torque relation observed in chronic stroke survivors.

In light of these uncertainties, the purpose of this study is to characterize the EMG-torque relations of calf muscles at different ankle joint angles on both paretic and non-paretic sides in chronic stroke survivors, and thus, to understand how the slope of such relations (i.e., muscle efficiency-related parameter) is altered by changing joint angles. We hypothesize that the slope becomes greater as calf muscle lengths shorten because the operating range of calf muscles is usually located in the ascending limb of active force-length curve, so that the maximum force gradually decreases with shortening muscle lengths (27–29). We also hypothesize that as the muscle length becomes shorter, a greater increase in the slope is more likely to be observed on the paretic side than on the non-paretic side. This is because the width of the active force-length curve may be narrower on the paretic side (19) and thus, the relative reduction in the magnitude of peak forces on the paretic side becomes greater at comparable muscle lengths.

As a secondary goal, we also seek to characterize the relation between the maximum joint torque at each joint angle and the EMG-torque slope at this joint angle. Our hypothesis here is that the greater the slope, the smaller the maximum joint torque. This is because the slope is a measure of muscular contraction efficiency.

In order to better understand the potential impact of intrinsic muscular changes of muscular contraction efficiency after, we propose to assess associations between the relative slope-related parameter and the relative shear wave speed (SWS) and between the relative slope-related parameter and the relative muscle thickness. SWS is a non-invasive measure of tissue stiffness (30), and it has been shown that shear waves travel faster in paretic MG muscles than in non-paretic muscles of chronic stroke survivors (31, 32).

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But here are the results.

Results

The total number of subjects examined was 9 for MG and LG analyses and 8 for SOL and ALL analyses. This is because we excluded the subject when reporting SOL and ALL as the EMG recording from the SOL muscle was not successful in one of our subjects.

Passive Range of Motion at Ankle Joint

The passive range of motion at ankle joint on the paretic side (32.0°; IQR = 30.0–40.0) was significantly smaller by ~25% than on the non-paretic side (42.0°, IQR = 37.3–45.0) (p = 0.031, d_z = 0.992). Although the median maximum PF (30.0°, IQR = 30.0–35.3) and DF (−3.0°, IQR = −6.3–1.5) on the paretic side was smaller than on the non-paretic side (PF: 34.0°, IQR = 30.0–37.8; DF: −5.0°, IQR = −11.3 to −3.8), there was no significant side-to-side difference in the maximum PF (p = 0.156, d_z = 0.548) or in the maximum DF (p = 0.063, d_z = 0.816).

Effect of Joint Angles on Maximum RMS EMG and Joint Torque

When compared to the non-paretic side, the maximum RMS EMG on the paretic side was significantly smaller in the MG (p = 0.003, η2p

= 0.683) and LG (p = 0.003, η2p = 0.685) muscles, but not in the TA (p = 0.122, η2p = 0.272) and SOL (p = 0.096, η2p = 0.307) muscles (Figure 3). However, the maximum RMS EMG was not a function of the ankle joint angles in all the muscles (TA: p = 0.104, η2p = 0.734; MG: p = 0.093, η2p = 0.747; LG: p = 0.366, η2p = 0.521; SOL: p = 0.172, η2p = 0.666). Moreover, there was no significant interaction between the side and the ankle joint angle in all the muscles (TA: p = 0.387, η2p = 0.507; MG: p = 0.271, η2p = 0.586; LG: p = 0.247, η2p = 0.604; SOL: p = 0.427, η2p = 0.481). Note that the magnitude of TA RMS EMG was considerably smaller than the other muscles, which indicates that the potential effects of muscle co-contraction would likely be negligible.

The overall magnitude of the maximum joint torque on the paretic side was significantly smaller than on the contralateral side (p = 0.001, η2p = 0.776; Figure 4), showing significant differences at all the angles (p < 0.01). Furthermore, the maximum joint torque increased as the ankle joint became dorsiflexed (p = 0.029, η2p = 0.845), showing that the maximum joint torque at the dorsiflexed angles beyond the neutral angles was significantly greater than at the plantarflexed positions (p < 0.05). There was a significant interaction between the side and the ankle joint angle (p = 0.049, η2p

= 0.808).

FIGURE 4

Figure 4. Maximum joint torque at each ankle angle. Negative angles indicate dorsiflexion. ^*a significant difference in pairwise comparisons between the side (p < 0.01). ^†a significant difference in torque between dorsiflexed angles and plantarflexed angles (p < 0.05).

EMG-Torque Relations

As the ankle angle became more plantarflexed, the slope of the EMG-torque relations increased progressively (Figure 5). Table 1 summarizes the coefficient a and b for each calf muscle and ALL. For the MG muscle, the coefficient a on the paretic side was significantly greater than on the non-paretic side (p = 0.008, d_z = 0.549), indicating the greater increase in the slope as the ankle joint becomes plantarflexed. There was no significant difference in the coefficient b between sides (p = 0.359, d_z = 0.089), indicating that the slope was not significantly different at the maximum DF angles.

FIGURE 5

Figure 5. Changes in EMG-torque slope as a function of ankle angle. Note that the effects of changing the ankle angle on the EMG-torque slope are more significant on the paretic side than on the non-paretic side. Markers and lines indicate raw data and fitted results from each individual (in different colors), respectively.

TABLE 1

Table 1. The coefficient a and b values for the slope of EMG-torque relations (S) and ankle plantarflexion angle (A). S = aA² + b.

Similar results were found in the other muscles. The coefficient a on the paretic side was significantly greater than on the non-paretic side (LG: p = 0.004, d_z = 0.548; SOL: p = 0.008, d_z = 0.802; ALL: p = 0.008, d_z = 0.601), whereas there was no significant difference in the coefficient b between sides (LG: p = 0.250, d_z = 0.501; SOL: p = 0.547, d_z = 0.357; ALL: p = 0.945, d_z = 0.315).

Torque-Slope Relations

For all the cases, the maximum voluntary isometric plantarflexion torque was smaller as the slope of the EMG-torque relations was greater (Figure 6). Table 2 summarizes the coefficient c for each calf muscle and ALL. The coefficient c was significantly smaller on the paretic side than on the non-paretic side (MG: p = 0.020, d_z = 1.190; SOL: p = 0.016, d_z = 1.326; ALL: p = 0.016, d_z = 1.321) except for LG (p = 0.098, d_z = 0.627). However, the coefficient c for pooled data from both sides was not significantly different with the coefficient c on the paretic side for all cases (MG: p = 1.000, d_z = 0.053; LG: p = 0.074, d_z = 0.622; SOL: p = 0.109, d_z = 0.666; ALL: p = 0.742, d_z = 0.040).

FIGURE 6

Figure 6. Relationship between maximum torque at each ankle angle and EMG-torque slope at the corresponding angle. Note that the greater slope, the smaller torque. The quicker decay in the torque is shown on the paretic side than on the non-paretic side, whereas there is no significant difference between the paretic and pooled data. Markers and lines indicate raw data and fitted results from each individual (in different colors), respectively.

TABLE 2

Table 2. The coefficient c values for normalized maximum plantarflexion torque at each joint angle (T) and slope of EMG-torque relations (S). T = cS⁻¹.

Correlation Analysis

A strong linear relationship was observed between the relative coefficient a and the relative maximum torque for MG (r = 0.783; p = 0.017; Figure 7A), SOL (r = 0.762; p = 0.037; Figure 7C), and ALL (r = 0.881; p = 0.007; Figure 7D). However, no significant relationship was found for LG (r = 0.617; p = 0.086; Figure 7B).

FIGURE 7

Figure 7. Relationship between the relative maximum torque and the relative coefficient a in medial gastrocnemius (A), lateral gastrocnemius (B), soleus (C), and the average of the muscles (D).

For the MG muscle, the relative coefficient a was smaller as the relative SWS measured at the neutral position was greater (r = −0.733; p = 0.031; Figure 8A). However, the relationship between the relative coefficient a and the relative muscle thickness was not significant (r = 0.017; p = 0.982; Figure 8B).

FIGURE 8

Figure 8. Relationship of the relative (non-paretic/paretic) coefficient a with the relative (paretic/non-paretic) shear wave speed (SWS) (A) or muscle thickness (MT) (B) measured in medial gastrocnemius.

Discussion

To summarize, the purpose of this study was to investigate: (1) the effect of changes in ankle joint angles on the muscular contraction efficiency of the calf muscles (i.e., the slope of EMG-torque relations); (2) the relation between the maximum joint torque at each joint angle and the EMG-torque slope at this angle; and (3) the association between the relative coefficient a and the relative muscle thickness or SWS. This is in order to understand the impact of intrinsic muscular changes on muscular contraction efficiency after stroke.

Our results show that the paretic side has a greater slope coefficient a (i.e., more increment in the slope as a function of ankle plantarflexion angle) and smaller c (i.e., steeper decay in the maximum torque as a function of the slope). There was also a strong linear relationship between the relative joint torque and the relative coefficient a, in the case of MG, SOL, or ALL. For the MG muscles, the relative coefficient a (i.e., muscular contraction efficiency) was negatively correlated with the relative SWS (i.e., muscle stiffness).